This invention relates generally to computed tomography imaging, and more particularly, to generating volumetric images using data collected from a digital x-ray panel.
In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the "imaging plane". The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a "view". A "scan" of the object comprises a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector. In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection technique. This process converts the attenuation measurements from a scan into integers called "CT numbers" or "Hounsfield units", which are used to control the brightness of a corresponding pixel on a cathode ray tube display.
Digital x-ray panels capable of sampling a 40 cm by 40 cm area in a single projection are known. With such panels, detector signals are sampled one row at a time by activating a trigger signal. The activation signals for row 1 to N are activated sequentially while all the cell signals for each row are read simultaneously. A typical cell size is 200 .mu.m.times.200 .mu.m. For a panel that covers a 40 cm.times.40 cm region, the resulting image size is roughly 2000.times.2000 pixels.
When the panel is positioned opposite an x-ray source and both the source and the panel rotate about the patient to collect projection data from different angles, volumetric CT data is obtained. Since all the x-rays diverge from the x-ray source to the detector in three dimensions, a set of cone shaped samples is obtained for each projection. This type of scanner is referred to as cone beam CT.
Cone beam reconstruction algorithms are typically required due to the divergence of the sampling rays in both x-y and z directions. A known cone beam reconstruction algorithm is described in Feldkamp et al., "Practical cone-beam algorithm," J. Opt. Soc. Am. A., vol. 1, no. 6, pp. 612-619. This algorithm pre-weights the projections according to a pre-defined weighting function. The weighted projection is then filtered and backprojected to generate the reconstructed images. Because the backprojection process is in a cone-beam geometry, a magnification dependent (therefore, location dependent) scaling factor is needed. The requirement of the scaling factor significantly increases the computational complexity of the algorithm, which makes the reconstruction process very slow.